1. Field of the Invention
This invention relates to an optical tomography method of obtaining an optical tomographic image by measurement of OCT (optical coherence tomography), and more particularly to an optical tomography method of obtaining an optical tomographic image by measurement of SD-OCT (spectral domain OCT).
2. Description of the Related Art
As a method of obtaining a tomographic image of an object of measurement such as living tissue, it is proposed to measure OCT (optical coherence tomography) as disclosed in Japanese Unexamined Patent Publication Nos. 6(1994)-165784 and 2003-139688. In the OCT measurement, a phenomenon that interference light is detected when the optical paths of the measuring light and the reflected light conform to the optical path of the reference light in length is used. That is, in this method, low coherent light emitted from a light source is divided into measuring light and reference light and the measuring light is projected onto the object of measurement, while the reflected light from the object of measurement is led to a multiplexing means. The reference light is led to the multiplexing means after its optical path length is changed in order to change the depth of measurement in the object. By the multiplexing means, the reflected light and the reference light are superposed one on another, and interference light due to the superposition is detected by, for instance, heterodyne detection.
In the above OCT system, a tomographic image is obtained by changing the optical path length of the reference light, thereby changing the measuring position (the depth of measurement) in the object. This technique is generally referred to as “TD-OCT (time domain OCT)”. More specifically, in the optical path length changing mechanism for the reference light disclosed in Japanese Unexamined Patent Publication No. 6(1994)-165784, an optical system which collects the reference light emitted from the optical fiber on a mirror is provided and the optical path length is adjusted by moving only the mirror in the direction of the beam axis of the reference light. Further, in the optical path length changing mechanism for the reference light disclosed in Japanese Unexamined Patent Publication No. 2003-139688, the reference light emitted from the optical fiber is turned to parallel light, the reference light in the parallel light is collected and caused to enter the optical fiber again by an optical path length adjusting lens, and the optical path length adjusting lens is moved back and forth in the direction of the beam axis of the reference light.
Whereas, as a system for rapidly obtaining a tomographic image without changing the optical path length of the reference light, there has been proposed an optical tomography system for obtaining an optical tomographic image by measurement of SD-OCT (spectral domain OCT). In the SD-OCT system, a tomographic image is formed without scanning in the direction of depth, by dividing broad band, low coherent light into measuring light and reference light by the use of a Michelson interferometer, projecting the measuring light onto the object and carrying out a Fourier analysis on each channeled spectrum obtained by decomposing the interference light of the reflected light, which returns at that time, and the reference light.
One of the characteristics of OCT is that the resolution in the direction of depth, that is, in the direction of the optical axis, is not limited by the NA (numerical aperture) of the optical system. That is, the resolution Δz in the direction of the optical axis is defined by the following formula (1), wherein λ and Δλ respectively represent the central wavelength of the measuring light and the full width half maximum of the spectrum of the measuring light.Δz=21n2/π(λ2/Δλ)  (1)
Accordingly, as shown in FIG. 4, even in a low NA optical system in which, for instance, λ=1.3 μm and Δλ=35 nm, a high resolution Δz (Δz=20 nm) in the direction of the optical axis can be realized.
The beam diameter Δx and the focal depth b of the measuring light beam in the converging position is defined by the following formulae (2) and (3), wherein the focal length and the diameter of the lens are respectively represented by f and d.Δx=4λ/π(f/d)  (2)b=πΔx2/2λ  (3)
FIG. 5 shows how the beam diameter changes with change of the position in the direction of the optical axis when the beam diameter Δx of the measuring light beam in the converging position which governs the lateral resolution of the tomographic image takes various values. In FIG. 5, curves 1, 2, 3, 4, 5 and 6 respectively show the cases where Δx=1 μm, 3 μm, 10 μm. 20 μm, 30 μm and 100 μm. Further, the position 0 in the direction of the optical axis represents the converging position. As shown in FIG. 5, when, for instance, Δx=20 μm, the bean diameter hardly changes even in a position about 750 μm away from the converging position in the direction of the optical axis, whereby Δx=20 μm is substantially held with the focal depth b within the range of b=1.5 mm.
The OCT systems are frequently for taking a tomographic image of an area of about 3 mm (in the direction of the optical axis)×5 mm (laterally: in the direction perpendicular to the optical axis) and initially have been developed for use in the field of ophthalmology. FIG. 6 shows a representative example of the TD-OCT systems. The TD-OCT system 200 comprises: a light source unit 110 comprising a light source 111 for emitting a light beam L and a collecting lens 112; a light dividing means 2 for dividing the light beam L emitted from the light source unit 110 to be propagated through an optical fiber FB1; a light dividing means 3 for dividing the light beam L passing therethrough into a measuring light beam L1 and a reference light beam L2; an optical path length adjusting means 20 for adjusting the optical path length of the reference light beam L2 divided by the light dividing means 3 to be propagated through an optical fiber FB3; an optical probe 30 that irradiates the measuring light beam L1 divided by the light dividing means 3 to be propagated through an optical fiber FB2 onto an object S of measurement; a multiplexing means 4 (the dividing means 3 doubles) for multiplexing a reflected light beam L3, which is the measuring light beam L1 reflected from the object S, and the reference light beam L2; and an interference light detecting means 240 for detecting interference light beam L4 of the reflected light beam L3 and the reference light beam L2 which have been multiplexed by the multiplexing means 4.
The optical path length adjusting means 20 comprises a collimator lens 21 which makes parallel the reference light beam L2 radiated from the optical fiber FB3, a mirror 23 which is movable in the direction of arrow A to change the distance to the collimator lens 21, and a mirror moving means 24 which moves the mirror 23 and changes the optical path length of the reference light beam L2, thereby changing the measuring position in the object S in the direction of depth. Further, on the optical path of the reference light beam L2 (in the optical fiber FB3), a phase modulating means 210 which slightly shifts the frequency of the reference light beam L2 is disposed. The reference light beam L2 which has been changed in its optical path length and shifted with its frequency by the optical path length adjusting means 20 is led to the multiplexing means 4.
The interference light detecting means 240 detects the intensity of the interference light L4 propagating through the optical fiber FB2 from the multiplexing means 4, for instance, by heterodyne detection. Specifically, when the sum of the whole optical path length of the measuring light beam L1 and the whole optical path length of the reflected light beam L3 is equal to the whole optical path length of the reference light beam L2, a beat signal which varies in intensity at the difference frequency between the reference light beam L2 and the reflected light beam L3 is generated. As optical path length is changed by the optical path length adjusting means 20, the measuring position (depth) in the object S is changed, whereby the interference light detecting means 240 comes to detect a plurality of beat signals which are generated in a plurality of the measuring position in the object S. The information on the measuring position is output to the image obtaining means from the interference light detecting means 240. Then an optical tomographic image is generated on the basis of the beat signals detected by the interference light detecting means 240 and information on the measuring position in the mirror moving means 24.
In the conventional OCT system having such a structure, the movement of the mirror 23 in the direction of the optical axis and the lateral scanning of the measuring light beam L1 are generally carried out at about one fourth of a desired resolution when an optical tomographic image is to be obtained. That is, when a resolution of about Δx=Δz=20 μm, the mirror 23 is moved in the direction of the optical axis at pitches of about 5 μm, and the light beam is laterally scanned at pitches of about 5 μm.
At present, a high resolution, high sensitivity and high speed OCT system is developed attempting application to a field other than the field of ophthalmology. More specifically, it is desired a high resolution system where the present resolution of about Δx=Δz=20 μm is improved to not higher than 10 μm and it is desired a super high resolution system where the present resolution of about Δx=Δz=20 μm is improved to not higher than 5 μm.
When Δx=Δz=10 μm, it will be understood from FIG. 5 that the distance in the direction of the optical axis over which Δx=10 μm is held is about 200 μm. Similarly, when Δx=Δz=5 μm, it will be understood that the distance in the direction of the optical axis over which Δx=5 μm is held is about 50 μm. Accordingly, it cannot be avoided a problem that, when a tomographic image of an area longer than a length in the direction of the optical axis where a desired beam diameter range can be held is intended to be obtained while a desired lateral resolution is kept to be held, the lateral resolution deteriorates in the deeper or shallower region.
As a method of overcoming the problem, a method disclosed, for instance, in “Ultrahigh-Resolution Optical Coherence Tomography by Broadband Continuum Generation from a Photonic Crystal Fiber”, Yimin Wang et al., OPTICS LETTERS, Vol. 28, No. 3, pp. 182-184, 2003, has been known. In this method, image data in the focusing position of the measuring light beam can be constantly obtained by changing the focusing position of the measuring light beam in synchronization with adjustment of the optical path length of the reference light beam and the method is called “dynamic focus OCT”.
However, the above dynamic focus OCT involves a problem that a long time is required to collect image data necessary for forming a tomographic image since the scanning mechanism for the measuring light beam must be fed at fine pitches in the direction of the optical axis. If so, though a sample removed from a living tissue can be measured, in vivo measurement on a living body while it moves is impossible.